Methods and computer program products for quantitative three-dimensional image correction and clinical parameter computation in optical coherence tomography

ABSTRACT

Methods and computer program products for quantitative three-dimensional (“3D”) image correction in optical coherence tomography. Using the methods and computer program products, index interface (refracting) surfaces from the raw optical coherence tomography (“OCT”) dataset from an OCT system can be segmented. Normal vectors or partial derivatives of the curvature at a refracting surface can be calculated to obtain a refracted image voxel. A new position of each desired refracted image voxel can be iteratively computed. New refracted corrected voxel positions to an even sampling grid can be interpolated to provide corrected image data. In some embodiments, clinical outputs from the corrected image data can be computed.

RELATED APPLICATION

The presently disclosed subject matter is a continuation of U.S. patentapplication Ser. No. 12/799,890, filed May 4, 2010, the disclosure ofwhich is incorporated herein by reference in its entirety.

GOVERNMENT INTEREST

This invention was made with government support under Grant Nos. R21EY017393 and K12-EY01633305 awarded by the National Institutes ofHealth. The government has certain rights in the invention.

TECHNICAL FIELD

The present subject matter generally relates to optical coherencetomography (“OCT”). More particularly, the present subject matterrelates to OCT for use in analysis of ocular optical elements andsurfaces, such as the cornea and lens of the eye.

BACKGROUND

Multiple modalities currently exist to image the ocular anterior segmentof an eye. These imaging devices aid in the clinical diagnosis and careof ocular disease. Optical coherence tomography (“OCT”) is one suchtechnique used to image in vivo the ocular anterior segment of an eye ina non-invasive fashion. Specific implementations can comprisetime-domain OCT systems (such as those sold under the name of VISANTE®by Carl Zeiss Meditec, Inc.), and Fourier-domain OCT systems, includingboth swept-source and spectrometer-based spectral-domain (“SDOCT”)implementations. Many previously disclosed anterior segment OCT systemshave utilized illumination light in the 1310 nm region. Recently,several SD-OCT systems have been developed that utilize an 840 nm lightsource and are designed for imaging the retina, but that can also bemodified with an adapter or a separate patient interface to image theocular anterior segment of an eye. These SD-OCT systems offer theadvantage of higher axial resolutions and faster scanning over priorsystems.

All OCT imaging systems are subject to the effects of refraction atsurfaces corresponding to interfaces between regions of differingrefractive index within a sample, including between the air and thesurface of the sample as well as internal sample interface surfaces. Fora sample such as the anterior segment of an eye, important refractiveindex interfaces comprise the outer, herein referred to as epithelial,and inner, herein referred to as endothelial surfaces of the cornea, aswell as the outer and inner surfaces of the crystalline lens.Additionally, for samples containing regions of different refractiveindex, it can be important for images acquired of the sample to reflectthe true physical dimensions of the sample rather than to be distortedby the varying speed of light in different sample regions. Both of thesepotential pitfalls can be particularly important in applications such ascorneal biometry where accurate measurements of various clinicallysignificant parameters must be computed from the image data. Suchcomputations are very sensitive to even small image errors due torefraction or distortion. Most current OCT systems do not correct theraw image data for refraction at sample interfaces or for the effects ofdifferent sample regions having differing refractive indices. Mostcurrent OCT systems instead assume that the light incident on the samplecontinues in a straight line through the sample and thus plot the rawimage A-scan data corresponding to a depth-resolved reflectivity map ofthe sample on this assumed undeviated path. These systems also do notcorrect for the effects of different refractive indices in differentregions of the sample. At best, these systems may divide the observedA-scan data by some assumed average index of refraction of the entiresample. As such, raw OCT data in current generation OCT systems does notaccurately represent the true position of internal sample structures andare thus not able to support calculation of clinically significantparameters which depend on accurate image data. In particular, toproduce accurate quantitative measurements of structures of the ocularanterior segment, accounting for the effects of refraction of the samplearm light and for the effects of differing indices of refraction indifferent sample regions is required.

Prior methods to correct for refraction in OCT images have beendeveloped. They do not, however, account accurately or completely forcorrection of volumetric, three-dimensional (“3D”) OCT datasets. Onemethod is limited to two-dimensional (“2D”) processing, which assumesthat refraction occurs only within the plane of individual acquiredB-scans (defined as sets of A-scans acquired along a straight linecomprising a cross-sectional image of the sample). For a curved 3Dstructure such as the cornea, 2D refraction correction is correct onlyif the sample is rotationally conically symmetric about some axispassing through the apex of the sample, and if the acquired B-scan datapassed exactly through the apex point. The first condition is rarelytrue for realistic samples such as the human cornea, especially if theyhave been modified surgically. The second condition is true only foridealized radial scan patterns, which may not be optimal because theyoversample the central region and undersample the outer region and inany case are difficult to obtain correctly due to unavoidable patientmotion or operator misalignment of the OCT system.

SUMMARY

In accordance with this disclosure, methods and tools for providingquantitative three-dimensional (“3D”) image correcting clinicalparameter computations for optical coherence tomography are provided. Itis, therefore, an object of the present disclosure to provide 3D imagingcorrection tools that can be used with any optical or ultrasonic imagingtools when curved surfaces are presented along the imaging path. Anotherobject of the present disclosure is to provide methods for improving the3D imaging of the eye to more accurately portray the refracting surfacesof the eye for diagnosis and treatment purposes.

This and other objects of the present disclosure as can become apparentfrom the present disclosure are achieved, at least in whole or in part,by the subject matter described herein.

BRIEF DESCRIPTION OF THE DRAWINGS

The patent or application file contains at least one drawing executed incolor. Copies of this patent or application with color drawing(s) willbe provided by the Patent and Trademark Office upon request and paymentof necessary fee.

A full and enabling disclosure of the present subject matter includingthe best mode thereof to one of ordinary skill in the art is set forthmore particularly in the remainder of the specification, includingreference to the accompanying figures, in which:

FIG. 1 illustrates a flowchart of an embodiment of a method forcorrection of refraction and index variations at one or more indexinterfaces in OCT images according to the present subject matter;

FIG. 2 illustrates an embodiment of a three-dimensional graph on aCartesian coordinate system of Snell's law for ocular anterior segmentOCT imaging according to the present subject matter;

FIG. 3 illustrates three-dimensional refraction correction methodparameters according to the present subject matter;

FIG. 4A illustrates an embodiment of a corneal scanning pattern in theform of a radial scanning pattern with each B-scan passing through theapex according to the present subject matter;

FIG. 4B illustrates an embodiment of a corneal scanning pattern in theform of a raster scanning pattern according to the present subjectmatter;

FIG. 5A illustrates an embodiment of refraction corrected voxels (e.g.,voxels on the endothelial surface) no longer on an even sampling gridafter a 3D Snell's law correction according to the present subjectmatter;

FIG. 5B illustrates an embodiment of a Zernike 3D interpolation withregenerated evenly sampled voxels according to the present subjectmatter;

FIG. 6A illustrates the accuracy of an embodiment of a Zernike 3Dinterpolation in which the mean Zernike 3D interpolation errors are lessthan 0.9 micrometers within a diameter 4-millimeter according to thepresent subject matter;

FIG. 6B illustrates the accuracy of an embodiment of a Zernike 3Dinterpolation in which the mean Zernike 3D interpolation errors are lessthan 2.8 micrometers for 7-millimeter diameter according to the presentsubject matter;

FIG. 7A illustrates an embodiment of the re-sampling issue in the radialscanning pattern according to the present subject matter;

FIG. 7B illustrates an embodiment of a re-sampling of a radial scanningpattern with a Zernike polynomial even-grid interpolation based on thenoneven-grid radial sampling points according to the present subjectmatter;

FIG. 8 illustrates an embodiment of residue errors from after 8^(th)order Zernike fitting of the refraction corrected inner surface of arigid contact lens according to the present subject matter;

FIG. 9 illustrates a top surface of an epithelium and a second surfaceof an endothelium according to the present subject matter;

FIG. 10A illustrates a 3D volume of raw corneal data where the DC linesform a horizontal plane according to the present subject matter;

FIG. 10B illustrates a 3D refraction corrected volumetric dataset wherethe corrected DC lines form a curved plane according to the presentsubject matter;

FIG. 10C illustrates a 3D corrected volume in another view angleaccording to the present subject matter;

FIG. 10D illustrates a normal thickness map of a subject's corneagenerated via the RHSA method according to the present subject matter;

FIG. 10E illustrates a thickness mapping of the cornea shown in FIGS.10A-10C obtained by a reference instrument (Oculus Pentacam);

FIG. 10F illustrates thickness mapping differences between 2D and 3DSnell's law according to the present subject matter;

FIG. 11 illustrates a computed thickness of a phantom rigid contact lenswith radial scanning pattern according to the present subject matter;

FIG. 12 illustrates a computed normal thickness map from spectral domainoptical coherence tomography data of an in vivo subject with the RHSAmethod according to the present subject manner;

FIG. 13A illustrates a diagrammatic representation of wavefrontaberration analysis according to the present subject matter;

FIG. 13B illustrates the results of wavefront aberration analysis usingZernike spectrum analysis and ray tracing according to the presentsubject matter;

FIG. 14A illustrates a float image (difference map of imaged surfaceagainst a reference surface) of the anterior surface of a contact lenswith asphericity fitting according to the present subject matter;

FIG. 14B illustrates a float image of the posterior surface of a contactlens with best sphere fitting according to the present subject matter;

FIG. 14C illustrates a float image of the posterior surface of a contactlens with asphericity fitting according to the present subject matter;and

FIG. 14D illustrates a float image of the posterior surface of a contactlens with best sphere fitting according to the present subject matter

FIG. 15 illustrates the cornea modeled as a thin lens and therelationship between the radii of curvature and the calculation ofrefractive power; and

FIG. 16 illustrates an float map (difference map of imaged surfaceagainst a reference surface—here a best fitting sphere) of the posteriorendothelial surface of an in vivo human cornea as derived from aSD-OCT-acquired volumetric dataset of a subject.

DETAILED DESCRIPTION

Reference will now be made in detail to possible embodiments of thepresent subject matter, one or more examples of which are shown in thefigures. Each example is provided to explain the subject matter and notas a limitation. In fact, features illustrated or described as part ofone embodiment can be used in another embodiment to yield still afurther embodiment. It is intended that the subject matter disclosed andenvisioned herein covers such modifications and variations.

For meaningful analysis of the cornea of an eye from spectral domainoptical coherence tomography (“SD-OCT”) data, effects of refraction atthe epithelial and endothelial surfaces should be accounted for. Amethod for three-dimensional (“3D”) refraction correction based on avector representation which accounts for refraction of optical coherencetomography (“OCT”) light in the cornea is present herein below. Thismethod can be used to dewarp raw SD-OCT volumes. The method can also beused to reconstruct the true position of the corneal endothelialsurface.

In addition, the raw volumetric image dataset may require correction ofimage artifacts or interpolation of non-sampled or under-sampled areasof the sample prior to further processing. Image artifacts can be causedby mis-calibrations or non-linearities of the image acquisition hardwareas well as patient-induced artifacts such as axial or lateral patientmotion, loss of patient fixation, or patient blinking during dataacquisition.

Following 3D refraction correction of volumetric corneal datasets, anestimate of the corneal thickness and the individual wavefrontaberrations of the epithelial and the refraction-corrected endothelialsurfaces can be made. To measure the corneal thickness along theepisurface normal vector, a numerical method or solution calledrecursive half searching algorithm (“RHSA”) can be used to compute thecorneal thickness with fine convergence. The different wavefrontaberrations of the anterior and posterior surfaces can be retrievedusing Zernike spectrum analysis. To enhance the scattering OCT signal atwide scanning range, the illumination light can be customized to aspecial polarization state to increase the scatting sensitivity. Ppolarization states are suggested to improve the imaging depth andreduce the corneal surface scattering. Different scanning patterns canalso be explored to optimize the OCT image quality and reduce bulkmotion. Asphericity and best spherical fitting can be utilized toretrieve the corneal asphericity, curvature and flow image of bothsurfaces of ocular anterior segment, which is very useful for clinicalapplications. The 3D refraction correction algorithms can be used withany optical or ultrasonic imaging tools when some curved surfaces arepresented along the imaging path.

Thus, a method is provided for three-dimensional (“3D”) correction ofrefraction and index variations within arbitrary samples. The method canbe particularly useful for correction of human ocular anterior segmentimages obtained in vivo using SDOCT systems including extraction ofquantitative clinical parameters from 3D OCT datasets acquired in thecornea. A mathematical model and analytical equations for correction ofrefraction and index variations can also be provided. Though this methodhas been reduced to practice for correcting OCT images of the ocularanterior segment and cornea, it can be applied generally to thecorrection of all imaging modalities whereby the source radiationencounters an interface and structure which alters its path (refractionof electromagnetic or sound waves).

Following correction of refraction and index variations, the correcteddata can then be used to generate multiple quantitative clinicallysignificant displays. These can comprise corneal thickness mapping,corneal surface curvature mapping, corneal surface deviation from bestfit sphere (float) mapping, corneal-iris angle mapping, and surface wavefront aberration analysis among others. Without 3D refraction correctionof the raw anatomic images, it would not be possible to generateaccurate (and hence clinically useful) versions of these data maps.These displays may be useful for evaluation of pathologies of the corneaor anterior segment, or planning of surgeries such as refractive surgery(e.g. LASIK) or intraocular lens surgery.

The subject matter described herein for quantitative 3D image correctionand clinical parameter computation in optical coherence tomography canbe implemented using a computer readable medium having stored thereonexecutable instructions that when executed by the processor of acomputer control the processor to perform steps. Exemplary computerreadable media suitable for implementing the subject matter describedherein includes disk memory devices, programmable logic devices, andapplication specific integrated circuits. In one implementation, thecomputer readable medium may comprise a memory accessible by aprocessor. The memory may comprise instructions executable by theprocessor for implementing any of the methods for quantitative 3D imagecorrection and clinical parameter computation in optical coherencetomography described herein. In addition, a computer readable mediumthat implements the subject matter described herein may be distributedacross multiple physical devices and/or computing platforms.

The following provides a description for correction of refraction andindex variations at one or more index interfaces in OCT images byapplication of Snell's law generalized to three-dimensions. FIG. 1illustrates a flowchart of an algorithm, generally designated 10, usedfor correction of refraction and index variations at one or more indexinterfaces in OCT images. To begin, a raw OCT dataset 12 can beprovided. A step 14 in the method 10 can comprise segmentation of indexinterface (refracting) surfaces from the raw OCT dataset 12. First, theexternal and internal sample surfaces at which refraction occurs can beidentified. For anterior segment imaging, these correspond to theanterior and posterior surfaces of the cornea and optionally of thecrystalline lens. Additionally, the locations within the sample at whichit is desired to correct the raw OCT image data for refraction and indexvariations is identified. These locations can correspond to allpositions below the first refracting surface or between all segmentedrefracting surfaces, or alternatively to only selected sample positionsat which it is known a priori that corrected data must be obtained inorder to calculate important derived clinical parameters.

For the case of anterior segment imaging, correction of refraction andindex variations for all image voxels, or volumetric pixels, cangenerate a complete de-warped 3D image suitable for viewing and analysiswith 3D visualization software tools such as those used in commonpractice in radiology for analysis of image data from computedtomography (CT) or magnetic resonance (MR) imaging. Alternatively, foranterior segment imaging, correction of refraction and index variationscan be limited to structures such as the location of the endothelialsurface of the cornea if all that is needed is computation of parametersrelated solely to the anterior and posterior corneal surfaces such asthe corneal thickness, corneal curvature, and corneal wavefrontaberrations.

Methods for segmentation of the refracting surfaces can comprise manualtechniques requiring human interaction such as marking several points onthe refracting surface followed by fitting of a polynomial or othersmooth function to the user-generated points. Various methods andsoftware algorithms for automatic segmentation of tissue internalstructures can also be used for segmenting of the refracting surfaces,including methods utilizing operations and/or mathematical techniquessuch as data thresholding, and use of image gradients, snakes, and graphcuts. Thus, in the first step, a raw OCT dataset 12 comprising imagebrightness as a function of x, y, and z coordinates can begenerated/received, and a segmented image dataset of the refractiveindex interface surfaces comprising sets of the x, y, and z coordinatesof the interface surface can be produced.

Another step 16 of the method and method 10 can be to calculate normalvectors or partial derivatives of the curvature at a first refractingsurface (or next refracting surface if iterating). To apply Snell's lawin 3D using the theoretical model disclosed herein, the normal vector ateach point on the segmented refracting surface where the OCT sample armlight strikes the sample can be calculated. If this is the firstiteration through the method or refraction, correction is only needed atthe first refracting surface, then this step can be applied to the firstrefracting surface (i.e. for anterior segment imaging, the epithelialcorneal surface). If this has already been done for the first refractingsurface, then it can be done for the next refracting surface.

The method for calculating the surface normal can depend upon the scanpattern utilized to obtain the raw 3D OCT dataset and upon thecoordinate system desired for conducting further computations. In thefollowing examples, a Cartesian coordinate system is used, although thepresent subject matter can be generalized to any other coordinate systemsuch as a polar coordinate system. For computations using the Cartesiancoordinate system, the partial derivatives of the surface curvature inthe x and y directions are needed as they relate to the surface normal,which can be calculated either from the surface normals or directly. Ifthe data was acquired using a raster scan pattern oriented in the x or ydirection, these derivatives can conveniently be calculated fromone-dimensional polynomial functions fit using the well-knownleast-squares fitting method and algorithm to each raster line in bothorthogonal directions. Alternatively, if the data was obtained using anyother non-raster scan pattern such as a radial scan pattern, the surfacepoints can first be fit to a 3D Zernike polynomial model following whichthe surface derivatives at every point may subsequently be calculatedanalytically or computationally.

Thus, in step 16, a segmented image dataset of the refractive indexinterface surfaces comprising sets of the x, y, and z coordinates of theinterface surface can be generated/received, and a set of surface normalvectors or partial derivatives of the curvature of a refractive surfacecan be produced. The refracting surface from which the normal vectors orpartial derivatives of its respective curvature are derived can be afirst refractive surface, or subsequent refractive surfaces, if aniteration is being performed.

After the step 16 of calculating normal vectors or partial derivativesof the curvature at a refracting surface, the method can comprise a step18 of iteratively computing the new position of each desired refractedimage voxel, or volumetric pixel. With knowledge of the direction vectorof the incident light at each position on the refracting surface andalso of the surface normal at each point, the new corrected position ofevery desired image voxel beneath that surface (down to the level of thenext refracting surface, if there is one) can be computed. For example,the new corrected position of each desired image voxel beneath thatsurface can be computed using Equations (10) disclosed hereinbelow. Thedirection vector of the incident light at each position can be knownfrom either a) for the first refracting surface, the direction fromwhich the sample arm light is incident on each position of this firstrefracting surface, which is a property of the OCT system which must bewell known and characterized, or b) for subsequent refracting surfaces,the direction at which the light was refracted from the previousrefracting surface.

The equations defined for this step can correct for both the effects ofrefraction at the sample interface and also for the new index ofrefraction of the medium following the interface. Thus, in step 18, araw OCT dataset comprising image brightness as a function of x, y, and zcoordinates, a segmented image dataset of the refractive index interfacesurfaces comprising sets of the x, y, and z coordinates of the interfacesurface, a set of surface normal vectors or partial derivatives of thecurvature of the first (next if iterating) refractive surface, and a setof incident light direction vectors can be generated/received, and a setof corrected x, y, and z coordinates of the new corrected position ofeach desired image voxel beneath that surface can be produced.

After step 18 of iteratively computing the new position of each desiredrefracted image voxel, the method can comprise step 20 of interpolatingnew refracted corrected voxel positions to an even sampling grid. Thecomputation of the new position of each refracted image voxel of theprevious step 18 can result in new position locations not located on theoriginal raw image data sampled grid and may thus be difficult to storeand to use in making further computations. To correct this, therefraction-corrected voxel reflectivity data can be re-sampled back tothe original Cartesian sampling grid. This re-sampling can beaccomplished using a number of well known re-sampling algorithms,including Gaussian curves, 3D Zernike curves and wavelets. In oneembodiment, step 20 can be performed by fitting the corrected voxel datato a 3D Zernike polynomial model, which can then be re-sampled at anydesired grid locations. Thus, in the step 20, a set of corrected x, y,and z coordinates of the new corrected position of every desired imagevoxel beneath that surface can be generated/received, and a set ofinterpolated corrected image brightness values as a function of x, y,and z coordinates of an even sampling grid can be produced.

In step 22, it can be determined whether more refracting surfaces areneeded. If so, the respective steps 16, 18, 20 of the method ofcalculating normal vectors or partial derivatives of the curvature atthe first (or next if iterating) refracting surface, iterativelycomputing the new position of each desired refracted image voxel, andinterpolation of new refracted corrected voxel positions to an evensampling grid can be repeated until all desired refracting surfaces havebeen evaluated enabling computation and resampling of all desired imagevoxels. If no more refracting surfaces are needed, then, at this point,a correct 3D dataset 24 can be considered to have been provided.

The next step 26 in the method 10 can comprise computing clinicaloutputs from corrected image data. From the three-dimensional refractionand index variation corrected voxels and surfaces, various clinicaloutputs can then be computed. Thus, in this step 26, a set ofinterpolated corrected image brightness values as a function of x, y,and z coordinates of an even sampling grid can be generated/received,and a set of clinical output measures, or parameters, 28 comprisingnumerical values, tables, graphs, images, and maps can be produced.

The method and algorithms described above are provided in more detailbelow.

Detailed Theory and Methods

In this section, the theoretical and computational details that can beused for implementation of the general method described in the previoussection are provided. The example application used for illustration ofthe method is a quantitative image correction and clinical parametercomputation for OCT analysis of the human cornea and anterior chamber,although it should be understood that the same inventive steps andprocedures can be applied to many other biological and industrial sampletypes.

The sample data and images used in this section to illustrate reductionto practice of the present subject matter were acquired using one of twoFourier-domain OCT systems. The first system is a research prototypeswept-source OCT system operating at 1310 nm wavelength. The secondsystem is a commercial high-speed SDOCT system operating at 840 nm witha corneal adapter manufactured by Bioptigen, Inc., of Durham, N.C. Thecommercial system can collect repeated A-scans from a 2048 pixel linescan camera at a readout rate of to 17 kHz. Imaging speed is not alimitation factor for any of the methods described. The axial resolutioncan be approximately 5 micrometers. For comparison, this is 10-foldgreater than the 50 micrometer resolution of a rotating Scheimpflugphotography system currently in clinical usage.

The goal of the system is a mathematical model of the human cornea andto use this system to generate equations that can accurately repositionrefracted pixels in an ocular anterior segment volumetric dataset.

1. Refraction Correction Theory and Algorithm

One method can be to merely use 2D refraction correction via the wellknown Snell's law. However, such a method requires an assumption thatthe corneal surface is rotationally conically symmetric which can bemathematically expressed as:

$\begin{matrix}{z = {z_{0} + \frac{{1/R} \times \left( {\left( {x - x_{0}} \right)^{2} + \left( {y - y_{0}} \right)^{2}} \right)}{1 + \sqrt{1 + {{K\left( {\left( {x - x_{0}} \right)^{2} + \left( {y - y_{0}} \right)^{2}} \right)}/R^{2}}}}}} & (1)\end{matrix}$

In the above expression, (x₀,y₀,z₀) represents the vertex of therotational conic; (x,y,z) is any point on the rotational conic; R is theapex radius of curvature of the surface; K is the conic parameter whichrepresents the asphericity of a surface. K=1 for a sphere and K=0 for aparabolic surface.

If this equation were true for the corneal surface, the cornea would beexactly rotationally symmetric, and 2D refraction correction would besufficient by carefully scanning each B-scan exactly through the axis ofrotation (apex). Unfortunately for clinical application, the cornea isgenerally not exactly rotationally symmetric. Therefore, the aboveassumption cannot be used. This is especially the case in pathologiccorneas with gross asymmetry and in surgically altered corneas (LASIK,corneal transplantation), and these corneas are the ones that canbenefit most from quantitative imaging.

To reflect physical reality, a general 3D vector implementation ofSnell's law can be applied to correct for refraction of any surfaceshape. The epithelium of the cornea can be mathematically implicitlydescribed as z_(EPI)−ƒ(x,y)=0. In the equation, the function ƒ(x, y) canbe expressed in terms of numerical partial derivatives of the measuredepithelium.

The derivation of Snell's law for use in three-dimensional applicationscan be done via Maxwell's equations or more intuitively. In theMaxwell's equations approach, the SDOCT instrument can be designed totelecentrically illuminate the cornea within an fixed diameter areacentered on the apex of the cornea, thus constraining the incident lightdirection across any scan pattern to the +z direction. In any realisticbiological or industrial sample, the magnetic current densities {rightarrow over (M)} are zero. The {right arrow over (n)} can be the unitsurface normal vector of an arbitrary incident point C on the epitheliumsurface. According to Maxwell's equations, the electric magnetic fieldboundary condition can be represented:{right arrow over (n)}×({right arrow over (E)} _(in) −{right arrow over(E)} _(ref))={right arrow over (M)}=0  (2)

Here {right arrow over (E)}_(in) and {right arrow over (E)}_(ref) arethe incident and refracted electric fields individually. The aboveequation indicates that the incident and refracted electric fields canbe continuous on the tangential surface of the incident epitheliuminterface. In other words, the refracted light still propagates withinthe same incident plane and only bends in that plane. This is true forboth p- and s-polarized incident light. In this regard, it can be usefulto point out that Brewster's angle for the cornea is Arctan(n_(c)/n_(air))≈54⁰, where n_(c) is the refractive index of cornea(1.385) and n_(air) is the refractive index of air (1.0). When thediameter of the scanning range is bigger than about 6 mm, the majorityof the incident light can be obliquely reflected by the epithelialsurface and thus lost if the incident polarization state is random.Thus, OCT images of the cornea in that region can be very weak. However,the incident angle can be quite close to Brewster's angle when thescanning range is bigger than 8 mm. As a result, part of the s-polarizedlight perpendicular to the incident plane will generally be reflected,but all the p-polarized light within the incident plane will generallybe transmitted with no loss. To enhance brightness of OCT images of thecornea for large scan widths, p-polarized light can be thus desirablealthough the cornea has strong birefringence effects which will changethe polarization states of the transmitted p-polarization light intoelliptical states.

FIG. 2 illustrates a 3D version of Snell's law for ocular anteriorsegment OCT imaging. In the graphed version, the following variables areprovided:

C(x₀,y₀,z₀): Incident point on the epithelium;

C′(x,y,z): intersection point of refraction light with beneath corneallayer,

{right arrow over (M)}C: incident telecentric light;

CC′: refraction light;

{right arrow over (CN)}∝{right arrow over (n)}: surface normal at pointC;

θ_(i): incident angle; and

θ_(r): refraction angle.

As illustrated in FIG. 2, the unit normal vector on the epithelialcorneal surface at point C can be expressed as:

$\begin{matrix}{\overset{\rightharpoonup}{n} = {{\overset{\rightarrow}{CN}/{{\overset{\rightarrow}{CN}/}}} = {\frac{\left( {{- \frac{\partial z_{EPI}}{\partial x}},{- \frac{\partial z_{EPI}}{\partial y}},1} \right)}{\sqrt{\left( \frac{\partial z_{EPI}}{\partial x} \right)^{2} + \left( \frac{\partial z_{EPI}}{\partial y} \right)^{2} + 1}}.}}} & (3)\end{matrix}$

The incident light {right arrow over (MC)} has direction vector (a,b,c),and the unit vector of the refracted ray {right arrow over (CC)}′ is:{right arrow over (CC)}′=(x−x ₀ ,y−y ₀ ,z−z ₀)/√{square root over ((x−x₀)²+(y−y ₀)²+(z−z ₀)²)}.  (4)Now (z−z₀) can be obtained using directional vector projection and itis:

$\begin{matrix}\begin{matrix}{\left( {z - z_{0}} \right) = {\sqrt{\left( {x - x_{0}} \right)^{2} + \left( {y - y_{0}} \right)^{2} + \left( {z - z_{0}} \right)^{2}} \cdot {\cos\left( {\theta_{in} - \theta_{r}} \right)}}} \\{= {{{OPL}/n_{c}} \cdot {\cos\left( {\theta_{in} - \theta_{r}} \right)}}}\end{matrix} & (5)\end{matrix}$

In Eq. (5), √{square root over ((x−x₀)²+(y−y₀)²+(z−z₀)²)} can bedirectly retrieved from the optical path length (“OPL”) as the distancefrom the refracting surface to the desired (x,y,z) image voxel in theraw OCT B-scan image. If the Eq. (2) is satisfied, the relationshipbetween the incident light and the refracted light can be represented bythe following 3D generalization of Snell's law:{right arrow over (MC)}×(−{right arrow over (n)})·n _(air) ={right arrowover (CC)}′×(−{right arrow over (n)})·n _(c).  (6)

Here × represents the vector cross product. In actual implementation ina Cartesian coordinate system, equation (6) actually can consist ofthree separate equations, one for each Cartesian axis.

For a more intuitive derivation of Eq. (6), Snell's law serves as thestarting point which describes the relationship between an incident rayof light and its refracted ray after encountering a refracting surface(in this case, the cornea):n _(air)*sin θ_(incident) =n _(cornea)*sin θ_(refracted).  (7)

Here, n_(air) is the refractive index of air; θ_(incident) is the anglebetween the incident ray of light and the surface normal; n_(comea) isthe refractive index of the cornea; and θ_(refracted) is the anglebetween the refracted ray of light and the continuation of the surfacenormal through the cornea.

The refraction occurs within the same plane. That is, the incident rayand the surface normal (the “incident plane”) exist in the same plane asthe refracted ray and the surface normal (the “refracted plane”). Theproblem with OCT is that this particular plane may not be shared withthe plane of the B-scan as is normally assumed with prior 2D refractioncorrection methods. Because the “incident plane” and “refracted plane”share the same larger plane, they also share the same normal vector tothis larger plane, particularly at the point of refraction. This sharednormal vector passing through the point of refraction can be termed asN. Both sides of Snell's law can be multiplied by this shared normalvector:n _(air)*sin θ_(incident) *

=n _(cornea)*sin θ_(refracted)*

.  (8)

By definition, for a vector

and another vector

, the cross product is a

×b

=|a

∥b

| sin [θ₁ab ]=N

where |a

| is the magnitude of vector {right arrow over (a)} and |

| is the magnitude of vector

; θ_(ab) is the angle between the two vectors; and

is the vector normal to the plane created by the two vectors. If the twovectors are unit vectors (each of magnitude 1), then the cross productof two vectors is simply the product of the sine of the angle betweenthe two vectors times the normal to the two vectors. Substituting thisdefinition into equation (8), a vector 3D form of Snell's law iscreated:n _(air) *

=n _(cornea)*(

×

).  (9)

Where

is the unit vector describing the incident ray;

is the unit vector describing the refracted ray; and

is the surface normal to the cornea for that particular incident ray.This is the same as equation (6) (where

is an alternative designation for

and

is an alternative designation for

).

Conversion of each vector to its Cartesian (x, y, z) form and solvingthe cross product in equation (6) assuming n_(air)=1.0 gives thefollowing equations for the refraction corrected positions in the x andy axes (z is given in equation (5) as derived by projection):

$\begin{matrix}\left\{ \begin{matrix}{{z - z_{0}} = {{{OPL}/n_{c}} \cdot {\cos\left( {\theta_{in} - \theta_{r}} \right)}}} \\{x = {\frac{a*{OPL}}{n_{c}^{2}} + x_{0} + {\frac{\partial z_{EPI}}{\partial x}\left( {{c*{{OPL}/n_{c}^{2}}} - z + z_{0}} \right)}}} \\{y = {\frac{b*{OPL}}{n_{c}^{2}} + y_{0} + {\frac{\partial z_{EPI}}{\partial y}\left( {{c*{{OPL}/n_{c}^{2}}} - z + z_{0}} \right)}}}\end{matrix} \right. & (10)\end{matrix}$

If the incident light is telecentrically scanned, then incident lightvector {right arrow over (MC)} becomes (0,0,−1) and equation (10) isfurther simplified as:

$\begin{matrix}\left\{ \begin{matrix}{{z - z_{0}} = {{{OPL}/n_{c}} \cdot {\cos\left( {\theta_{in} - \theta_{r}} \right)}}} \\{x = {x_{0} - {\frac{\partial z_{EPI}}{\partial x}\left( {{{OPL}/n_{c}^{2}} + z - z_{0}} \right)}}} \\{y = {y_{0} - {\frac{\partial z_{EPI}}{\partial y}\left( {{{OPL}/n_{c}^{2}} + z - z_{0}} \right)}}}\end{matrix} \right. & (11)\end{matrix}$

FIG. 3 provides illustrative definitions of the parameters in Equations(10-11). In particular, FIG. 3 illustrates the 3D refraction correctionmethod parameters, which can be as follows:

θ_(in): Incidental angle;

θ_(r): Refraction angle;

OPL: Optical path length in OCT image; and

z: The projection of OPL on the z-axis.

The 3D refraction corrected volume can then be generated by applyingEquations (10-11) to each desired voxel subsequent to the refractingsurface in a 3D volumetric dataset. Equations (10-11) can enablepractical 3D refraction and index of refraction variation correction ofvolumetric OCT datasets.

Many scanning patterns can be employed in ophthalmic imaging. One commonscanning pattern can be raster scanning (a stack of sequential B-scansis used to form a box of images) and another common scanning pattern canbe radial scanning (each B-scan is centered around a rotational axis).The advantage of the raster scanning is that it provides complete andevenly sampled x, y, and z data throughout the volume. For example, eachB-scan exists in the x-axis, the stacking of B-scans creates the y-axis,and the individual points in an A-scan define the z axis. This is idealfor Cartesian implementation of the 3D refraction correction. While theradial scanning pattern has the advantage of high SNR for all B-scanframes, the data is generally asymmetrically sampled with densersampling centrally and less peripherally. For a Cartesian implementationas in equation (10), this requires that the data first be interpolatedand then resampled evenly. One such implementation is via Zernikefitting as described below. Alternatively, a polar implementation ofequation (6) can be created.

Examples of these two different scanning patterns are illustrated inFIGS. 4A and 4B. In particular, FIG. 4A illustrates a radial scanningpattern with each B-scan passing through the apex. FIG. 4B illustrates araster scanning pattern.

Although these are two of the more common scanning patterns, it shouldbe understood that scanning can comprise other patterns of B-scansdesigned for optimal coverage of the area of interest in the anteriorsegment. Alternatively, the dataset may comprise multiple individualA-scans distributed around the eye's surface in such a way as tomaximize the concentration of image data in regions of particularclinical interest or in such a way as to minimize image artifacts suchas caused by motion of the patient during data acquisition. Suchpatterns of individual A-scans can place them in radial, circular,raster, or random patterns, or patterns that contain repeated or nearlyrepeated sequences of sub-patterns designed to minimize the effects ofpatient motion.

2. Zernike 3D Interpolations

2.1. Zernike 3D Interpolations for Raster Scanning Pattern

One issue with the newly generated, 3D refraction corrected volumetricdata can be that the pixels are no longer evenly sampled. As such, ameans to evenly resample such data for further analysis andvisualization can be devised. For example, the complete endothelialsurface can be reconstructed to obtain a corneal thickness map (eachepithelial normal may no longer intersect a physical endothelial pixel)or to generate wavefront aberration analyses of the endothelium. Toreconstruct even-sampled epithelium and the endothelium surfaces, aZernike polynomial 3D interpolation can be employed to obtain theeven-grid epithelial, uncorrected endothelial and refraction correctedendothelial surfaces. By using Zernike interpolation to reconstruct thesurfaces, image re-segmentation from the newly generated refractioncorrected volumetric dataset can be avoided.

Zernike polynomials are a set of complete orthogonal polynomials definedon a unit circle. These polynomials are represented in polar coordinatesby:

$\begin{matrix}{{Z_{i}\left( {r,\theta} \right)} = {{R_{n}^{m}(r)}{\Theta^{m}(\theta)}}} & (12) \\{{R_{n}^{m}(r)} = {\sum\limits_{s = 0}^{{({n - {m}})}/2}\frac{\left( {- 1} \right)^{s}\sqrt{n + 1}{\left( {n - s} \right)!}r^{n - {2s}}}{{{{{s!}\left\lbrack {{\left( {n + m} \right)/2} - s} \right\rbrack}!}\left\lbrack {{\left( {n - m} \right)/2} - s} \right\rbrack}!}}} & (13) \\{{\Theta^{m}(\theta)} = \left\{ \begin{matrix}{\sqrt{2}\cos{m}\theta} & \left( {m > 0} \right) \\1 & \left( {m = 0} \right) \\{\sqrt{2}\sin{m}\theta} & \left( {m < 0} \right)\end{matrix} \right.} & (14) \\{{A\left( {{Rr},\theta} \right)} = {\sum\limits_{i = 0}^{\infty}{c_{i}{Z_{i}\left( {r,\theta} \right)}}}} & (15) \\{c_{i} = {\frac{1}{\pi}{\int_{0}^{2\pi}{\int_{0}^{1}{{A\left( {{Rr},\theta} \right)}{Z_{i}\left( {r,\theta} \right)}r\ {\mathbb{d}r}\ {\mathbb{d}\theta}}}}}} & (16)\end{matrix}$

Where the indices n and m are the radial degree and the azimuthalfrequency respectively; i is a mode-ordering number; R is used tonormalize the measurement data to a unit circle and c_(i) is Zernikeelevation interpolation coefficients.

FIGS. 5A and 5B show the procedure for Zernike 3D interpolation. In FIG.5A, the volumetric pixels, or voxels, from the 3D refraction correctedvolumetric data are not evenly sampled. In particular, the refractioncorrected voxels (e.g., voxels on the endothelial surface) are no longeron an even sampling grid after 3D Snell's law correction. Through theuse of the Zernike polynomial 3D interpolation, the voxels from the 3Drefraction corrected volumetric data can be evenly sampled again asshown in FIG. 5B. Thus, any elevation value of arbitrary points can beobtained using Zernike interpolation coefficients.

Since OCT can employ a very dense scanning pattern containing more thanten thousand points, the 3D Zernike interpolation can provide a goodinterpolation of the corneal surface without loss of accuracy. Forinstance, using a 16 order Zernike polynomial fit, the mean fitting(residual) error can be 0.9 micrometers within a 4-millimeter diameterand 2.8 micrometers for a 7-millimeter diameter. FIGS. 6A and 6B showthe Zernike 3D interpolation errors. In FIG. 6a , the mean Zernike 3Dinterpolation errors are less than 0.9 micrometers within a diameter4-millimeter. In FIG. 6B, the mean interpolation errors are less than2.8 micrometers for 7-millimeter diameter. Both interpolation errors arewithin the range of the light source coherence length of 6.4micrometers.

Since these errors are much less than the coherence length of 6.4micrometer, the Zernike 3D interpolation is acceptable for clinicalapplications. By using Zernike coefficients in representing cornealsurfaces also can provide the following advantages.

-   -   (1) For large data volumes, only Zernike coefficients (less than        200 float points) can be used to significantly compress the data        scale.    -   (2) The coefficients only represent the particular surface but        not the other layers. Hence, image re-segmentation is        unnecessary.    -   (3) The coefficients provide the elevation of any arbitrary        point, which intrinsically avoids the additional interpolation        steps when trying to find the intersection points between        anterior surface normals and the posterior surface.

It is not necessary to do the 3D interpolation for the epithelialsurface in the raster scanning pattern as it is already evenly sampled.

2.2. Zernike 3D Interpolations for Radial Scan Patterns

Similarly, Zernike 3D interpolation can be performed for all surfaces innon-raster scan patterns (such as radial scan patterns). In a radialscan pattern, the central area of the scan can be sampled more denselythan that in the periphery. FIGS. 7A and 7B illustrate the samplingdensity in radial scanning cases. FIG. 7A shows a radial scanningpattern. As can be seen, the sampling density is different betweencentral and peripheral points. Specifically, the sampling points aremuch sparser in the outer circles than in the inner circles.

FIG. 7B shows a Zernike polynomial even-grid interpolation performed fora surface that has been sampled radially for both the epithelium andendothelium surfaces based on the noneven-grid radial sampling points.This applies to all scan patterns which sample unevenly (with respect tox and y).

The interpolation procedure for this case is summarized as following:

-   (1) Segment both the epithelium and endothelium surface.-   (2) Employ an 8th-order Zernike polynomial fitting to interpolate    both surfaces into an even-grid sampling space.-   (3) Calculate the surface partial differential equations of the    epithelial surface.-   (4) Use equation (10) to obtain the refraction corrected endothelium    surface.-   (5) Utilize the Zernike 3D interpolation again to reconstruct the    refraction corrected endothelial surface.

FIG. 8 shows an example of 8th-order Zernike fitting residue errors ofrefraction corrected endothelium surface of a rigid contact lens. Themaximum fitting errors can be less than 0.7 micrometers. While the useof Zernike functions represents a particularly attractive solution tore-sampling of the optically refracted data points to an evenly sampledgrid, it should be understood that any number of other re-samplingtechniques which are well known in the art may also be used for thispurpose.

3. Other Data Correction

Instrument mis-calibrations and non-linearities can be rectified bycareful calibration of the instrument on a calibration target of knowndimensions prior to patient imaging, including calculation of correctionfactors to be applied to the resulting image data. Correction forpatient motion or loss of fixation can be rectified by employing methodsto estimate the actual patient head or eye motion which occurred duringimage acquisition, and then using this information to correct theresulting image data. Methods to estimate the actual patient motion cancomprise external instrumentation such as an iris camera or eye tracker(both used to estimate lateral or rotational eye motion), or cancomprise image processing methods such as image registration to estimateaxial or lateral motion in OCT A-scans, B-scans, or volume scans.

For example, patient axial or lateral motion between B-scans can beestimated from the peak of the normalized cross-correlation betweensequential A-scans (for axial motion), B-scans (for axial and lateralmotion in the direction of the B-scan), or volume scans (for axial andlateral motion, assuming the volume scans are acquired quickly comparedto the patient motion). Many other image registration and motionestimation algorithms are well known in the art of image processing andcan be used for this purpose. Once estimates of axial or lateral motionare obtained, they can be used to correct the volumetric image datasetby re-positioning image data within the dataset to its true position.Also, if the image dataset contains missing image data (e.g., becausethe patient blinked or because lateral motion resulted in some areas ofthe sample not being undersampled), such data can be filled in by imageinterpolation algorithms. Other bulk motion methods can includeanalyzing the natural motion frequency of the patient. Then the samplingfrequency can be adjusted to any high order harmonic frequency of thepatient motion. In this way, it is possible to statistically reduce thepatient motion.

Correction of image artifacts can be done as just described upon the rawA-scan, B-scan, or volumetric image datasets, creating corrected imagedatasets before further segmentation and image processing.Alternatively, rather than creating entirely new corrected imagedatasets, it can be possible and computationally simpler to perform someartifact correction operations upon partially processed data, such assegmented data.

Correction of some image artifacts, such as those resulting from axialor lateral patient motion or missing data, can be accomplished at thispoint when the entire image dataset has been reduced to a mathematicalsurfaces model. For example, a simple but effective method forcorrection of axial patient motion during sequential radial B-scanacquisition comprises registering the segmented corneal anterior surfaceof each radial B-scan to that of a selected reference B-scan (e.g., thefirst one in the set) at a selected lateral location within each B-scan(e.g., its center pixel). This axial motion estimate can then be appliedto correct the entire mathematical surface model such that the center ofeach radial B-scan is at the same axial location. This method isexpected to work well when the axial component dominates patient motionduring data acquisition.

4. Clinical Parameter Computation: Corneal Thickness Mapping, WaveAberration Analysis, Asphericity Analysis, Refractive Curvature Maps,Float Maps, and Others

4.1 Corneal Thickness Mapping

A corneal thickness map can be created by determining the normaldistance between the epithelial and endothelial corneal surfaces in thevolumetric surface model via an analytical or numerical approach. Asused herein, the term normal can be defined as normal to the epithelialsurface. There can be a thickness value for each pixel representing theepithelial surface. The thickness value for each pixel can be mapped toa color scale to produce a two-dimensional pachymetry (thickness) map.

The corneal thickness is measured along the epithelial surface normalvector. The surface normal equation of the epithelium is obtained by:

$\begin{matrix}{\frac{\left( {x - x_{0}} \right)}{{\partial{F\left( {x,y,z} \right)}}/{\partial x}} = {\frac{\left( {y - z_{0}} \right)}{{\partial{F\left( {x,y,z} \right)}}/{\partial y}} = \frac{\left( {z - z_{0}} \right)}{- 1}}} & (17)\end{matrix}$

Where (x₀,y₀,z₀) stands for the coordinate of point C; and (x,y,z)stands for the coordinate of an arbitrary point on the normal equation(line).

It can be difficult to derive an analytical solution between equation(17) provided above and equation (18) provided below. A numericalsolution called the recursive half searching algorithm (“RHSA”) andmethod can be used here to solve for (x, y, z) that exists within adefined distance (e.g., an error tolerance) to the endothelial surface.RHSA starts with an estimate of where the endothelial pixel should bealong the normal line that was calculated as part of the 3D refractioncorrection. If that estimate does not exist on the prior Zernikeinterpolation of the endothelial surface, it will be off by some amountin the z direction. That amount is halved and used as the next estimateof the endothelial pixel. The process is repeated until the differencebetween the estimate and the surface is less than the desired error.

For instance, if the error tolerance is 0, then (x, y, z) will becoincident with the endothelial surface. An appropriate error toleranceis chosen to balance clinical needs and computation time. The principleis illustrated in FIG. 9 of searching intersection points between thesurface normal equation of epithelial and refraction-correctedendothelial surface. In FIG. 9, the top surface is epithelium and thesecond is endothelium. The syllables illustrated in FIG. 9 are asfollows:

z′: Obtained using equation (16) after half distance recursion; and

z″: retrieved using Zernike coefficients in equation (15).

Thus, RHSA permits searching for the normal (perpendicular) distancebetween two surfaces given known spatial coordinates of the firstsurface and given a representation of the second surface in a basisspace. In this particular implementation, the basis functions areZernike polynomials, and the RHSA permits searching for the intersectionbetween the epithelial surface on the normal equation and theendothelium.

The starting guess point can be the refraction uncorrected endothelium(x,y,z). The new coordinates of (x,y,z) of the first step can beestimated by manipulating equation (17):

$\begin{matrix}{{Step}\mspace{14mu} 1\text{:}\left\{ \begin{matrix}{z^{\prime} = {z\left( {{uncorrected}\mspace{14mu}{endothelium}\mspace{14mu}{position}} \right)}} \\{x^{\prime} = {x_{0} - {{\partial{F\left( {x,y,z} \right)}}/{\partial{x_{({{x\; 0},{y\; 0},{z\; 0}})}\left( {z^{\prime} - z_{0\;}} \right)}}}}} \\{y^{\prime} = {y_{0} - {{\partial{F\left( {x,y,z} \right)}}/{\partial{y_{({{x\; 0},{y\; 0},{z\; 0}})}\left( {z^{\prime} - z_{0\;}} \right)}}}}}\end{matrix} \right.} & (18)\end{matrix}$

∂F(x,y,z)/∂x_((x0,y0,z0)) and ∂F(x,y,z)/∂y_((x0,y0,z0)) can be theconstant partial derivatives of the epithelial surface at point C, andthey do not need to be calculate again. In step 2, it can be determinedif the new point (x′, y′, z′) is located at the Zernike interpolatedendothelial surface.

$\begin{matrix}{{Step}\mspace{14mu} 2\text{:}\left\{ \begin{matrix}{z^{''} = {\overset{\infty}{\sum\limits_{i = 0}}{c_{i}{Z_{i}\left( {{r\left( {x^{\prime},y^{\prime},z^{\prime}} \right)},{\theta\left( {x^{\prime},y^{\prime},z^{\prime}} \right)}} \right)}}}} \\{{{{{if}\mspace{14mu}{{z^{\prime} - z^{''}}}} \leq ɛ},{{then}\mspace{14mu}{stop}\mspace{14mu}{and}\mspace{14mu}\left( {x^{\prime},y^{\prime},z^{\prime}} \right)\mspace{14mu}{will}\mspace{14mu}{be}\mspace{14mu}{the}\mspace{11mu}{target}\mspace{14mu}{point}}}\;}\end{matrix} \right.} & (19)\end{matrix}$

ε is the stop condition (i.e., the error tolerance), which can bedefined as 1 micrometer in some examples. Step 3 will be followed if itdoes not reach the convergent accuracy.

$\begin{matrix}{{Step}\mspace{14mu} 3\text{:}\left\{ \begin{matrix}{{{if}\mspace{14mu}{{z^{\prime} - z^{''}}}} > {ɛ\mspace{20mu}{then}}} \\{z^{\prime} = {{z^{\prime} + {{{{z^{\prime} - z^{''}}}/2}\mspace{14mu}{if}\mspace{14mu} z^{\prime}}} < {z^{''}\mspace{14mu}{searchpointisbelowendothelium}}}} \\{{z^{\prime} = {{z^{\prime} - {{{{z^{\prime} - z^{''}}}/2}\mspace{14mu}{if}\mspace{14mu} z^{\prime}}} > {z^{''}\mspace{14mu}{searchpointisaboveendothelium}}}}\;} \\{{RepeatStep}\; 1\mspace{14mu}{and}\mspace{14mu}{Step}\; 2}\end{matrix} \right.} & (20)\end{matrix}$

The above search method can be used to recursively halve the differenceof the search distance. Hence, it can converge very fast and only needs6 steps to define the intersection point.

Once each epithelial point in the modeled surface has a thickness valuecalculated for it, the values can be fit to a color scale. A typicalcolor scale can represent average corneal values (around 500-600 μm) asgreen with thinner values trending warmer (red) and thicker valuestrending cooler (blue).

The data can then be plotted as a two-dimensional map providing an enface (looking coincident with the direction of the visual axis) viewcorneal map. This concept of thickness mapping can be applied to any twodesired surfaces, for instance the thickness of the human lens can bemeasured between the front and posterior lens surfaces as imaged bySDOCT. This could also be applied to the magnitude of spaces between twosurfaces. For instance, the anterior chamber depth can be defined as thedistance between the corneal endothelial surface and the anterior lenssurface (or iris in the periphery or an anterior chamber lens ifpresent).

FIGS. 10A-10F illustrate example 3D refraction correction results. FIG.10A illustrates a raster scanned 3D raw corneal volumetric datasetwithout refraction correction. As illustrated, all DC lines (i.e.,artifacts in SDOCT images which are perpendicular to the axis ofillumination of the sample) form a horizontal plane. After 3D refractiondewarping, the corneal volume can be shrunken in the x, y and zdimensions simultaneously as shown in FIG. 10B. As shown in FIG. 10B,the horizontal plane formed by DC becomes curved after the refractioncorrection. FIG. 10C shows a 3D visualization from a different viewingangle. FIG. 10D shows the thickness mapping via RHSA method. The OculusPentacam measurement result of the same cornea is displayed in FIG. 10E(Pentacam is an accepted commercial instrument which measures cornealthickness using a different method than OCT). In this illustration, thedifference between the OCT and Pentacam measurement at the apex is 8micrometers which is within the measurement accuracy (50 micrometers) ofPentacam. FIG. 10F indicates the difference in thickness mapping between2D and 3D refraction correction. The mean difference between 2D and 3DSnell's law correction is 34 micrometers with a standard deviation of 30micrometers. The mean difference is large enough to be of clinicalsignificance. For example, the corneal thickness map was smoothed by a2-dimensional low-pass image filter. FIG. 11 shows the thickness map ofa phantom rigid gas permeable contact lens using a radial scanningpattern. FIG. 12 shows a thickness (pachymetry) map of an in vivo humancornea derived from spectral domain optical coherence tomographyacquired volumetric dataset of a subject.

Another method for determining the normal distance between theepithelial and endothelial corneal surfaces can be to use ray tracingand Newton's algorithm to find the intersection point along the Newtonsearch direction.

$\begin{matrix}{{{Newton}\mspace{14mu}{Search}\mspace{14mu}{Direction}} = {\quad\left\lbrack {{\frac{\partial{G\left( {x,y,z} \right)}}{\partial x}l},{\frac{\partial{G\left( {x,y,z} \right)}}{\partial y}m},{\frac{\partial{G\left( {x,y,z} \right)}}{\partial z}n}} \right\rbrack}} & (21)\end{matrix}$

G(x,y,z)=0 is the implicit form of the endothelium and the directioncosines of the epithelial surface normal equation is represented in(l.m.n) which is related to the epithelial surface derivatives. Thismethod requires iterative calculation of all the partial derivatives,

$\quad\left\lbrack {\frac{\partial{G\left( {x,y,z} \right)}}{\partial x},\frac{\partial{G\left( {x,y,z} \right)}}{\partial y},\frac{\partial{G\left( {x,y,z} \right)}}{\partial z}} \right\rbrack$of the intermediate approximation points on the endothelium. Thus, thismethod can be considered computationally more intensive. In addition,the intermediate partial derivatives of the endothelial surface are muchmore sensitive to the noise which influences the convergence.

Still other methods which are well known in the art may be applied fordetermining the normal distance between the epithelial and endothelialcorneal surfaces following 3D refraction correction.

4.2 Corneal Wavefront Aberration Analysis

After 3D refraction correction, the individual wavefront aberration ofthe epithelium and endothelium layers can be estimated. Similar to themethod employed in professional optical design software Zemax™, theoptical path differences, also called wavefront aberrations W_(a),between the off-axis and on-axis for the all meridians are calculatedwithin 7-millimeters. Equation (22) gives the computing method.

$\begin{matrix}{W_{a} = {{n_{Cornea} \cdot f} - {n_{Air} \cdot Z_{a}} - {n_{Cornea} \cdot d}}} & (22) \\{C_{i} = {\frac{1}{\pi}{\int_{0}^{2\pi}{\int_{0}^{1}{{W_{a}\left( {{Rr},\theta} \right)}{Z_{i\;}\left( {r,\theta} \right)}r\ {\mathbb{d}r}\ {\mathbb{d}\theta}}}}}} & (23)\end{matrix}$

C_(i) describes the wavefront aberration coefficients. For example, C₁₃(with mode index=1) is called spherical aberration error. The wavefrontaberration analysis using Zernike spectrum analysis and ray tracing isillustrated in FIGS. 13A and 13B. In FIG. 13A, the anterior surfacecorneal height is labeled Z_(a); f is the distance of the focus point(F) to the apex of the surface VA; and d is the distance of this focuspoint to an arbitrary point on the surface. FIG. 13B shows the Zernikeaberration analysis of the cornea. Red is for anterior surface and greenis for posterior surface. Overall the anterior surface contributes muchmore aberration errors than that of posterior surface. The sphericalaberration error of the epithelium is around 1.1 micrometers and −0.6micromeres for the endothelium.

4.3 Asphericity, Refractive Curvature, and Best Fit Sphere Analysis

Asphericity and best sphere fitting can be used to estimate theasphericity, surface curvature, and to create float maps.

Asphericity (K) and the radius of curvature for a surface (R) can beobtained using 3D fitting with Equation (1). Using a phantom contactlens, the measurement error of the base curvature from that reported forthe phantom was less than 2.9%. The radius of curvature can be used todescribe the refractive power of the cornea either over an opticalregion (for instance, refractive power within the central 3 mm opticalzone) or locally over the entire cornea to produce a local refractivecurvature map.

Best sphere fitting can be retrieved using 3D fitting by simply lettingK=1 in equation (1). Examples of the results are demonstrated in FIGS.14A-14D. Float images of different 3D fitting of a phantom contact lensare shown therein. FIG. 14A shows a float image of the anterior surfaceof a contact lens with asphericity fitting. In FIG. 14A, the asphericityis 0.8 and radius is 10.03 mm. FIG. 14B shows a float image of theposterior surface of a contact lens with best sphere fitting. FIG. 14Cshows a float image of the posterior surface of a contact lens withasphericity fitting. In FIG. 14C, the asphericity is 0.59 and radius is9.32 mm. FIG. 14D shows a float image of the posterior surface of acontact lens with best sphere fitting. The residuals demonstrate theamount of deviation of the corneal surface from that of the best fitsphere. This can be a clinically important indicator of ecstatic cornealconditions.

Elevation (float) maps can represent the distance between the modeledsurface and a reference surface. For example, a sphere can serve as thereference surface, although any surface could be used (e.g., ellipsoids,other toric surfaces, defined planes). In the case of a referencesphere, the sphere can be chosen to best fit the modeled surface withina defined area of that surface. Corneal elevation maps can be created byfirst designating a reference surface. The reference shape can be chosenas a standard shape (e.g., a planar or spherical surface) or othersurface shape chosen to fit the modeled corneal surface. The deviationin the z-axis (nominally the same direction as the visual axis) of eachpixel in the modeled corneal surface from the reference surface can bedesignated as the elevation value. The pixel elevation values can againbe mapped to a color scale to produce this two-dimensional map.

Returning again to Equation (1), by setting K to 1, the surface ismodeled as spherical. Other values deform this surface into variousellipsoids. For this particular implementation, (x₀, y₀) can be assignedto either the center of the dataset or allowed to “float” such thatthere are no constraints on the location of this apex. The (x, y, z)values of the modeled surface within the area of desired fitting canthen be used to solve for R (and x₀, y₀, z₀ if unconstrained) using acurve fitting algorithm (e.g., least squares).

With K set to 1, this solution represents the best fitting sphere forthe modeled surface within the desired area and the residuals from thefitting represent the deviation in the z-axis of the modeled surface ateach (x, y, z) from the best fitting sphere. These residuals can be fitto a color scale and plotted to provide a two-dimensional en face viewof the elevation data, such as is shown in FIG. 16. This type ofreferenced elevation map can be created for any desired surface derivedfrom OCT images (corneal epithelial surface, endothelial surface,lenticular surfaces, etc.).

Regarding curvature and power maps, previous methods for calculatingcurvature values were based on estimates of radii of curvature of thecorneal epithelial surface using reflection based imaging technologies.With OCT, however, curvature values of any surface or combination ofsurfaces can be produced. For instance, a corneal curvature map can becreated by determining the radii of curvature at each local point(defined by some local radius, typically 0.1-1 mm) on the modeledsurfaces. Radii of curvature of the modeled corneal epithelial surfacecan be determined from the best fitting sphere as in elevation mappingfor small discrete, defined regions on the surface (e.g., regions thatare 1 mm in diameter). The radii of curvature can then be converted intodioptric power by convention using the following equation:

$\begin{matrix}{{Power} = \frac{n_{keratomtric} - 1}{r_{c}}} & (24)\end{matrix}$Here, r_(c) is the radius of curvature in mm within the discrete areaand n_(keratometric) can either be an accepted value (such as one toequate a r_(c) of 7.5 mm with 45 D) or empirically derived to yieldother dioptric power to r_(c) associations.

There are multiple alternative ways to constrain the local surfacefitting. The tangential curvature method can allow for the best fittingsphere/curve for the sampled surface area. The axial curvature methodcan constrain the origin of the radius of curvature of the best fittingsphere/curve to the visual axis. In this case, for equation (1), theradius of curvature would be explicitly stated to always originate at(x₀, y₀).

With an evenly spaced sampling of the radii of curvature across theepithelial surface, the values can be fit to a color scale and plottedto provide a two-dimensional en face representation of curvature withinterpolation as required if the sampling is insufficiently dense.Multiple color scales are available in the literature including thosecreated by standards bodies.

The radii of curvature may then be converted to ophthalmic diopterpowers, which can be inversely related to the radii according to theformulas in FIG. 15. This can be done using curvature data from each ofthe two corneal surfaces. Alternatively, the so-called “keratometric”dioptric power of the cornea can be estimated from the radius ofcurvature of the epithelial surface only by assumption of a so-calledkeratometric index of refraction n_(k). As shown in FIG. 15, R_(a) andR_(p) are the radii of the corneal anterior and posterior surfaces,which are often assumed to be spherical. The relevant refractive indicesare n₀=1.0 (air), n₁=1.376 (corneal stroma), n₂=1.336 (aqueous humor),and n_(k)=1.3375 (keratometric index). The curvatures and ophthalmic(refractive) powers thus calculated for each pixel or local region onthe cornea can then be mapped to a color scale to producetwo-dimensional representations of curvature and power, respectively.

For the specific case in which the surface area of the cornea to befitted is centered on the visual axis, this power can provide a summaryestimate of corneal optical power through the thin lens equation. Thisarea can typically be around 3-4 mm in diameter, as this is the regionof the cornea which is most spherical and thus most amenable to thespherical model described by radii of curvature. This calculationrequires derivation of the radius of curvature of both the anterior andposterior surfaces.

The above analyses (thickness, elevations and curvatures, estimatedoptical power) can also be performed for OCT images of the human lens.Because OCT provides a full volumetric rendering of both surfaces of thecornea, independent radii of curvature of both the anterior andposterior surfaces centered on any arbitrary axis can be calculated. Ifthese axes are sampled evenly across the cornea, a full map of physical(in contrast to keratometric) power of defined sampled areas could becreated similar to the maps described above (fit to color scale anddisplay en face).

Further, OCT images that contain both the cornea and lens (eitherregistered or in whole) can be used to create power maps describing thetwo refractive components of the human eye. These can be summed withappropriate optical treatment via ray tracing to determine the powercontributions of the cornea and lens in the visual axis or any arbitrarydesired axis. Unique, novel parameters such as percentage contributionof the corneal power (i.e., lenticular power) to the total power (i.e.,cornea power plus lens power) can be calculated and fit to a color scalesuch that 50% is centered on green with warmer colors representinggreater contribution of that component versus cooler colors representinglesser contribution of that component. This analysis can provideinformation for surgeons planning selective refractive procedures (e.g.,cornea or lens based).

As the cornea and lens comprise the main refractive components of theeye, analysis of both components from OCT images (singly and incombination as in power maps centered on designated axes as in thevisual axis) can provide valuable information regarding the refractivestatus of the eye. For volumetric OCT images of the anterior segmentthat comprise the irido-corneal angle, a similar map can be created ofthe magnitude of the angle or any numerical parameter derived fromimages of the irido-corneal angle (such as a specific defined 2-D areaof the angle within the image, distance of the iris to the cornea at aspecified distance from the apex of the angle).

4.4 Other Clinical Analyses

Given an anatomically accurate, 3D refraction corrected representationof the eye, other analyses can be carried out.

For instance, the angle between the iris and the adjacent cornealendothelial surface can be an important parameter in glaucoma care. Anaccurate volumetric dataset can be used to create an en-face map of themagnitude of these angles. Multiple radial samplings of the intersectionof the iris and adjacent corneal endothelial surfaces can be used toderive vectors from which the angle between them can be calculated.Similarly, the angle could also be described as any numerical quantitywhich describes the relevant clinical concept, such as the distancebetween the iris and cornea in this region and thus the gross physicalavailability of the trabecular meshwork. For instance, the triangulararea with its apex at the apex of the angle and its base defined as theline normal to a fixed reference such as Schwalbe's line or simply thelength of the base of this triangle. These samplings can then be mappedto their enface position. In contrast to current isolated local anglemeasurements, such an “angle map” can show regions of narrow angles orangle closure, which could be important to glaucoma management. The onlycurrent comparable clinical view would be direct manual visualization ofthe angle structures using a mirror and mentally (via drawing)recreating the observed degree of regional angle magnitudes. This typeof en-face angle magnitude map display from tomographic image data hasnot been previously described.

3D displays of the anterior chamber space can also be created. Thesedimensions can be useful in planning for artificial lens placements bothin the anterior chamber and posterior chamber.

If the OCT dataset includes data from the lens, the lenticular surfacescan also be serially refraction corrected using equation (9) atdifferent locations. For example, the lenticular surfaces can beserially refraction corrected first at the corneal epithelium, then atthe corneal endothelium, and then at the anterior lenticular surface.From the corrected data, wavefront aberration analyses of the lenticularsurfaces as described above can be undertaken.

As provided above, methods and algorithms are provided for quantitativeimage correction and clinical parameter computation which is generallyapplicable for any OCT sample containing refracting interfaces andregions of different refractive indices. The methods and algorithms areparticularly suitable for quantitative correction of 3D OCT images ofthe cornea and anterior segment of the eye. Two specific implementationsfor two different scanning patterns can be introduced for cornealimaging. Zernike 3D interpolation can be used to represent the cornealsurfaces (epithelium, uncorrected endothelium, and refraction correctedsurfaces). This interpolation method can make it possible for theimplementation of a recursive half searching algorithm (RHSA) and methodto measure the corneal thicknesses and map them in an en-face clinicalview. 3D corneal volumetric refraction correction can provide thefoundation for generating further clinical parameters. These compriseknown clinical ones such as wavefront analysis, asphericity, refractivecurvature maps, and best fit sphere float maps as well as novel onessuch as angle magnitude maps. 3D refraction correction and the accuraterepresentation of ocular structures it creates provide an important toolin the visualization and management of ocular disease.

Embodiments of the present disclosure shown in the drawings anddescribed above are exemplary of numerous embodiments that can be madewithin the scope of the appended claims. It is contemplated that theconfigurations of methods for quantitative three-dimensional imagecorrection and clinical parameter computation in optical coherencetomography can comprise numerous configurations other than thosespecifically disclosed.

What is claimed is:
 1. A method for quantitative three-dimensional(“3D”) image correction of optical coherence tomography (OCT) using acomputer readable medium having stored thereon executable instructions,that when executed by a processor of a computer control the computer toperform the method comprising: calibrating a scanning system of an OCTimaging system by imaging a calibration target with known physicaldimensions including deriving numerical correction factors that describenonlinearities of the scanning system; obtaining a three-dimensional OCTimage of a subject with the OCT imaging system; segmenting at least afirst refracting surface interface from an OCT image dataset for the OCTimage; correcting voxel positions for at least the segmented firstrefracting surface interface using the numerical correction factorsderived during calibration; and calculating a refracted image voxelposition for at least a portion of voxels in the OCT image dataset. 2.The method according to claim 1, further comprising repeating the stepsof segmenting a refracting surface interface and calculating a refractedimage voxel position until all desired refracting surfaces have beenevaluated.
 3. The method according to claim 1, wherein segmentingcomprises identifying external and internal sample surfaces at whichrefraction occurs.
 4. The method according to claim 3, wherein thesample surfaces comprise anterior and posterior surfaces of a cornea ora lens.
 5. The method according to claim 3, wherein the step ofsegmenting comprises identifying locations within the sample at which tocorrect the OCT image dataset for refraction and index variations. 6.The method according to claim 5, wherein the identified locationscomprise positions being at least one of below the first refractingsurface or between all segmented refracting surfaces.
 7. The methodaccording to claim 5, wherein the identified locations comprise selectedsample positions at which it is known that corrected data is to beobtained in order to calculate clinical parameters.
 8. The methodaccording to claim 1, wherein calculating refracted image voxelpositions comprises calculating a normal vector or partial derivativesof curvature at each point on the segmented refracting surface where anOCT sample arm light strikes the sample.
 9. The method according toclaim 8, wherein calculating normal vectors or partial derivatives ofthe curvature at the refracting surface is dependent upon a scan patternutilized to obtain the OCT image dataset and upon a coordinate systemdesired for conducting further computations.
 10. The method according toclaim 9, wherein the coordinate system is at least one of a Cartesiancoordinate system or a polar coordinate system.
 11. The method accordingto claim 10, wherein the coordinate system is a Cartesian coordinatesystem, and calculating normal vectors or partial derivatives of thecurvature at the refracting surface comprises calculating the partialderivatives of the surface curvature in the x and y directions fromsurface normals.
 12. The method according to claim 11, furthercomprising using a raster scan pattern oriented in the x or y directionto acquire data and calculating derivatives from one-dimensionalpolynomial functions fit using a least-squares fitting algorithm to eachraster line in both orthogonal directions.
 13. The method according toclaim 11, further comprising using a radial scan pattern to acquire dataand fitting surface points to a 3D Zernike polynomial model andsubsequently calculating the surface derivatives at every point by atleast one of analytically or computationally methods.
 14. The methodaccording to claim 11, wherein calculating normal vectors or partialderivatives of the curvature at the refracting surface comprises usingan equation of:{right arrow over (MC)}×(−{right arrow over (n)})·n _(air) ={right arrowover (CC)}′×(−{right arrow over (n)})·n _(c) where {right arrow over(MC)} comprises a direction vector (a,b,c) of incident light; ×represents a vector cross product; {right arrow over (n)} is a unitsurface normal vector of an arbitrary incident point C; n_(air) isrefractive index of air; {right arrow over (CC)}′ is unit vector of arefracted ray; and n_(c) is refractive index of subject and wherecalculations are applied after correcting the OCT image dataset forscanning non-linearities using the correction factors derived duringcalibration.
 15. The method according to claim 14, wherein calculatingnormal vectors or partial derivatives of the curvature at the refractingsurface comprises using equations of: $\quad\left\{ \begin{matrix}{{z - z_{0}} = {{{OPL}/n_{c}} \cdot {\cos\left( {\theta_{in} - \theta_{r}} \right)}}} \\{x = {\frac{a*{OPL}}{n_{c}^{2}} + x_{0} + {\frac{\partial z_{EPI}}{\partial x}\left( {{c*{{OPL}/n_{c}^{2}}} - z + z_{0}} \right)}}} \\{y = {\frac{b*{OPL}}{n_{c}^{2}} + y_{0} + {\frac{\partial z_{EPI}}{\partial y}\left( {{c*{{OPL}/n_{c}^{2}}} - z + z_{0}} \right)}}}\end{matrix} \right.$ to calculate refraction corrected positions in x-,y-, and z-axes, where C(x₀,y₀, z₀) is an incident point on anepithelium; θ_(in) is an incidental angle; θ_(r) is a refraction angle;OPL is an optical path length in an OCT image; and z is a projection ofOPL on the z-axis and where calculations are applied after correctingthe OCT image dataset for scanning non-linearities using the correctionfactors derived during calibration.
 16. The method according to claim 1,wherein calculating the position of each desired refracted image voxelcomprises generating a raw OCT image dataset comprising image brightnessas a function of x, y, and z coordinates, a segmented image dataset ofthe index interface refracting surfaces comprising sets of the x, y, andz coordinates of the refracting surface, a set of surface normal vectorsor partial derivatives of curvature of the refractive surfaces, and aset of incident light direction vectors and where calculations areapplied after correcting the OCT image dataset for scanningnon-linearities using the correction factors derived during calibration.17. The method according to claim 16, wherein calculating the positionof each desired refracted image voxel comprises producing a set ofcorrected x, y, and z coordinates of the corrected position of eachdesired image voxel beneath the refracting surface and wherecalculations are applied after correcting the OCT image dataset forscanning non-linearities using the correction factors derived duringcalibration.
 18. The method according to claim 1, further comprisinginterpolating of new refracted corrected voxel positions to an evensampling grid comprising re-sampling refracted corrected voxelreflectivity data back to an original Cartesian sampling grid.
 19. Themethod according to claim 18, wherein interpolating of new refractedcorrected voxel positions to an even sampling grid comprises fittingcorrected voxel data to a 3D Zernike polynomial model for re-sampling toa desired grid location.
 20. The method according to claim 18, whereininterpolating of new refracted corrected voxel positions to an evensampling grid comprises generating a set of corrected x, y, and zcoordinates of the new corrected position of each desired image voxelbeneath the refracting surface and where calculations are applied aftercorrecting the OCT image dataset for scanning non-linearities using thecorrection factors derived during calibraton.
 21. The method accordingto claim 20, wherein interpolating of new refracted corrected voxelpositions to an even sampling grid further comprises producing a set ofinterpolated corrected image brightness values as a function of x, y,and z coordinates of the even sampling grid.
 22. The method according toclaim 1, further comprising computing clinical outputs from correctedimage data.
 23. The method according to claim 22, wherein computingclinical outputs from corrected image data comprises producing a set ofclinical output measures comprising numerical values, tables, graphs,images, and maps.
 24. The method according to claim 22, wherein the OCTimage dataset is of a cornea or a lens, and wherein computing clinicaloutputs comprises one or more of generating thickness maps, generatingelevation maps, generating curvature maps, computing wavefrontaberrations, or computing an angle between an iris and an endothelialsurface.
 25. The method according to claim 22, wherein the OCT imagedataset is of a cornea and wherein computing clinical outputs fromcorrected image data comprises creating a corneal or lens thickness mapby determining normal distance between the two surfaces in thevolumetric surface model via an analytical or numerical approach. 26.The method according to claim 25, wherein a Cartesian coordinate systemis used to conduct further computations and a recursive half searchingalgorithm (“RHSA”) numerical solution is employed to solve within apredefined tolerance for perpendicular distances between discrete pointsof the epithelial surface and an interpolated second surface.
 27. Amethod for quantitative three-dimensional (“3D”) image correction inoptical coherence tomography using a computer readable medium havingstored thereon executable instructions, that when executed by aprocessor of a computer, control the computer to perform the methodcomprising: calibrating a scanning system of an OCT imaging system byimaging a calibration target with known physical dimensions includingderiving numerical correction factors that describe the nonlinearitiesof the scanning system; obtaining a 3D image of a subject with the OCTimaging system; segmenting of index interface refracting surfaces from araw OCT image dataset from an OCT system; calculating normal vectors orpartial derivatives of a curvature at a refracting surface to obtain arefracted image voxel; iteratively computing a new position of eachdesired refracted image voxel; interpolating of new refracted correctedvoxel positions to an even sampling grid to provide corrected imagedata; and correcting voxel positions using the numerical correctionfactors derived during calibration.
 28. The method according to claim27, further comprising repeating the steps of calculating normal vectorsor partial derivatives of a curvature at a refracting surface to obtaina refracted image voxel, iteratively computing a new position of eachdesired refracted image voxel, and interpolating of new refractedcorrected voxel positions to an even sampling grid to provide correctedimage data until all desired refracting surfaces have been evaluatedenabling computation and resampling of all desired image voxels.
 29. Themethod according to claim 27, wherein segmenting comprises identifyingexternal and internal sample surfaces at which refraction occurs. 30.The method according to claim 29, wherein the sample surfaces compriseanterior and posterior surfaces of a cornea or a lens.
 31. The methodaccording to claim 29, wherein segmenting comprises identifyinglocations within the sample surfaces at which to correct the raw OCTimage dataset for refraction and index variations.
 32. The methodaccording to claim 31, wherein the identified locations comprisepositions being at least one of below the first refracting surface orbetween all segmented refracting surfaces.
 33. The method according toclaim 31, wherein the identified locations comprise selected samplepositions at which it is known that corrected data is to be obtained inorder to calculate clinical parameters.
 34. The method according toclaim 27, wherein calculating normal vectors or partial derivatives ofthe curvature at the refracting surface comprises calculating a normalvector at each point on the segmented refracting surface where an OCTsample arm light strikes a sample.
 35. The method according to claim 27,wherein calculating normal vectors or partial derivatives of thecurvature at the refracting surface is dependent upon a scan patternutilized to obtain the raw OCT dataset and upon a coordinate systemdesired for conducting further computations.
 36. The method according toclaim 35, wherein the coordinate system is at least one of a Cartesiancoordinate system or a polar coordinate system.
 37. The method accordingto claim 36, wherein the coordinate system is a Cartesian coordinatesystem, and wherein calculating normal vectors or partial derivatives ofthe curvature at the refracting surface comprises calculating thepartial derivatives of the surface curvature in the x and y directionsfrom surface normals.
 38. The method according to claim 37, furthercomprising using a raster scan pattern oriented in the x or y directionto acquire data and calculating derivatives from one-dimensionalpolynomial functions fit using a least-squares fitting algorithm to eachraster line in both orthogonal directions.
 39. The method according toclaim 37, further comprising using a radial scan pattern to acquire dataand fitting surface points to a 3D Zernike polynomial model andsubsequently calculating surface derivatives at every point by at leastone of analytically or computationally.
 40. The method according toclaim 37, wherein calculating normal vectors or partial derivatives ofthe curvature at the refracting surface comprises using an equation of:{right arrow over (MC)}×(−{right arrow over (n)})·n _(air) ={right arrowover (CC)}′×(−{right arrow over (n)})·n _(c) where {right arrow over(MC)} comprises a direction vector (a,b,c) of incident light; ×represents a vector cross product; {right arrow over (n)} is a unitsurface normal vector of an arbitrary incident point C; n_(air) isrefractive index of air; {right arrow over (CC)}′ is unit vector of arefracted ray; and n_(c) is refractive index of cornea.
 41. The methodaccording to claim 40, wherein the coordinate system is a Cartesiancoordinate system, and wherein calculating normal vectors or partialderivatives of the curvature at the refracting surface comprises usingequations of: $\quad\left\{ \begin{matrix}{{z - z_{0}} = {{{OPL}/n_{c}} \cdot {\cos\left( {\theta_{in} - \theta_{r}} \right)}}} \\{x = {\frac{a*{OPL}}{n_{c}^{2}} + x_{0} + {\frac{\partial z_{EPI}}{\partial x}\left( {{c*{{OPL}/n_{c}^{2}}} - z + z_{0}} \right)}}} \\{y = {\frac{b*{OPL}}{n_{c}^{2}} + y_{0} + {\frac{\partial z_{EPI}}{\partial y}\left( {{c*{{OPL}/n_{c}^{2}}} - z + z_{0}} \right)}}}\end{matrix} \right.$ to convert each vector to its Cartesian form andsolving the cross product of the equation: {right arrow over(MC)}×(−{right arrow over (n)})·n_(air)={right arrow over(CC)}′×(−{right arrow over (n)})·n_(c), assuming n_(air)=1.0, whereC(x₀, y₀, z₀) is an incident point on an epithelium; θ_(in) is anincidental angle; θ_(r) is a refraction angle; OPL is an optical pathlength in an OCT image; and z is a projection of OPL on the z-axis. 42.The method according to claim 27, wherein the step of interpolating ofnew refracted corrected voxel positions to an even sampling gridcomprises re-sampling refracted corrected voxel reflectivity data backto an original Cartesian sampling grid.
 43. The method according toclaim 27, wherein the step of interpolating of new refracted correctedvoxel positions to an even sampling grid comprises fitting correctedvoxel data to a 3D Zernike polynomial model for re-sampling to a desiredgrid location.
 44. The method according to claim 27, wherein the step ofinterpolating of new refracted corrected voxel positions to an evensampling grid comprises generating a set of corrected x, y, and zcoordinates of the new corrected position of each desired image voxelbeneath the refracting surface.
 45. The method according to claim 44,wherein the step of interpolating of new refracted corrected voxelpositions to an even sampling grid further comprises producing a set ofinterpolated corrected image brightness values as a function of x, y,and z coordinates of the even sampling grid.
 46. A non-transitorycomputer readable medium having stored thereon executable instructions,that when executed by a processor of a computer, control the computer toperform steps comprising: calibrating a scanning system of an opticalcoherence tomography OCT imaging system by imaging a calibration targetwith known physical dimensions including deriving numerical correctionfactors that describe the nonlinearities of the scanning system;segmenting of index interface refracting surfaces from a raw OCT imagedataset from an OCT system; correcting voxel positions for at least thesegmented first refracting surface using the numerical correctionfactors derived during calibration; calculating a refracted image voxelposition for at least a portion of voxels in the raw OCT image dataset;and computing clinical outputs from a corrected image data.
 47. Thenon-transitory computer readable medium according to claim 46,comprising repeating the step of calculating a position of each desiredrefracted image voxel, and further comprising interpolating newrefracted corrected voxel positions to an even sampling grid to providecorrected image data until all desired refracting surfaces have beenevaluated enabling computation and resampling of all desired imagevoxels.
 48. The non-transitory computer readable medium according toclaim 46, wherein segmenting of index interface refracting surfaces fromthe raw OCT image dataset from an OCT system comprises identifyingexternal and internal sample surfaces at which refraction occurs. 49.The non-transitory computer readable medium according to claim 48,wherein the sample surfaces comprise anterior and posterior surfaces ofa cornea or a lens.
 50. The non-transitory computer readable mediumaccording to claim 48, wherein segmenting comprises identifyinglocations within the sample surfaces at which to correct the raw OCTimage dataset for refraction and index variations.
 51. Thenon-transitory computer readable medium according to claim 50, whereinthe identified locations comprise positions being at least one of belowthe first refracting surface or between all segmented refractingsurfaces.
 52. The non-transitory computer readable medium according toclaim 50, wherein the identified locations comprise selected samplepositions at which it is known that corrected data is to be obtained inorder to calculate clinical parameters.
 53. The non-transitory computerreadable medium according to claim 46, wherein calculating refractedimage voxel positions comprises calculating a normal vector or partialderivatives of the curvature at each point on the segmented refractingsurface where an OCT sample arm light strikes the sample.
 54. Thenon-transitory computer readable medium according to claim 53, whereincalculating normal vectors or partial derivatives of the curvature atthe refracting surface is dependent upon a scan pattern utilized toobtain the raw OCT image dataset and upon a coordinate system desiredfor conducting further computations.
 55. The non-transitory computerreadable medium according to claim 54, wherein the coordinate system isat least one of a Cartesian coordinate system or a polar coordinatesystem.
 56. The non-transitory computer readable medium according toclaim 55, wherein the coordinate system is a Cartesian coordinatesystem, and wherein calculating normal vectors or partial derivatives ofthe curvature at the refracting surface comprises calculating thepartial derivatives of the surface curvature in the x and y directionsfrom the surface normals.
 57. The non-transitory computer readablemedium according to claim 56, further comprising using a raster scanpattern oriented in the x or y direction to acquire data and calculatingthe derivatives from one-dimensional polynomial functions fit using aleast-squares fitting algorithm to each raster line in both orthogonaldirections.
 58. The non-transitory computer readable medium according toclaim 56, further comprising using a radial scan pattern to acquire dataand fitting surface points to a 3D Zernike polynomial model andsubsequently calculating surface derivatives at every point by at leastone of analytically or computationally.
 59. The non-transitory computerreadable medium according to claim 53, wherein calculating normalvectors or partial derivatives of the curvature at the refractingsurface comprises using an equation of:{right arrow over (MC)}×(−{right arrow over (n)})·n _(air) ={right arrowover (CC)}′×(−{right arrow over (n)})·n _(c) where {right arrow over(MC)} comprises a direction vector (a,b,c) of incident light; ×represents a vector cross product; {right arrow over (n)} is a unitsurface normal vector of an arbitrary incident point C; n_(air) isrefractive index of air; {right arrow over (CC)}′ is unit vector of arefracted ray; and n_(c) is refractive index of subject and wherecalculations are applied after correcting the raw OCT image dataset forscanning non-linearities using the correction factors derived duringcalibration.
 60. The non-transitory computer readable medium accordingto claim 59, wherein a Cartesian coordinate system is used, and whereincalculating normal vectors or partial derivatives of the curvature atthe refracting surface comprises using equations of:$\quad\left\{ \begin{matrix}{{z - z_{0}} = {{{OPL}/n_{c}} \cdot {\cos\left( {\theta_{in} - \theta_{r}} \right)}}} \\{x = {\frac{a*{OPL}}{n_{c}^{2}} + x_{0} + {\frac{\partial z_{EPI}}{\partial x}\left( {{c*{{OPL}/n_{c}^{2}}} - z + z_{0}} \right)}}} \\{y = {\frac{b*{OPL}}{n_{c}^{2}} + y_{0} + {\frac{\partial z_{EPI}}{\partial y}\left( {{c*{{OPL}/n_{c}^{2}}} - z + z_{0}} \right)}}}\end{matrix} \right.$ to convert each vector to its Cartesian form andsolving a cross product of the equation: {right arrow over(MC)}×(−{right arrow over (n)})·n_(air)={right arrow over(CC)}′×(−{right arrow over (n)})·n_(c), assuming n_(air)=1.0, whereC(x₀, y₀, z₀) is an incident point on an epithelium; θ_(in) is anincidental angle; θ_(r) is a refraction angle; OPL is an optical pathlength in an OCT image; and z is a projection of OPL on the z-axis andwhere calculations are applied after correcting the raw OCT imagedataset for scanning non-linearities using the correction factorsderived during calibration.
 61. The non-transitory computer readablemedium according to claim 46, wherein calculating the position of eachdesired refracted image voxel comprises generating the raw OCT imagedataset comprising image brightness as a function of x, y, and zcoordinates, a segmented image dataset of the index interface refractingsurfaces comprising sets of the x, y, and z coordinates of therefracting surface, a set of surface normal vectors or partialderivatives of the curvature of the refractive surfaces, and a set ofincident light direction vectors and where calculations are appliedafter correcting the raw OCT image dataset for scanning non-linearitiesusing the correction factors derived during calibration.
 62. Thenon-transitory computer readable medium according to claim 61, whereincalculating the position of each desired refracted image voxel comprisesproducing a set of corrected x, y, and z coordinates of the correctedposition of each desired image voxel beneath the refracting surface andwhere calculations are applied after correcting the raw OCT imagedataset for scanning non-linearities using the correction factorsderived during calibration.
 63. The non-transitory computer readablemedium according to claim 47, wherein interpolating of new refractedcorrected voxel positions to an even sampling grid comprises re-samplingrefracted corrected voxel reflectivity data back to an originalCartesian sampling grid.
 64. The non-transitory computer readable mediumaccording to claim 47, wherein interpolating of new refracted correctedvoxel positions to an even sampling grid comprises fitting correctedvoxel data to a 3D Zernike polynomial model for re-sampling to a desiredgrid location.
 65. The non-transitory computer readable medium accordingto claim 47, wherein interpolating of new refracted corrected voxelpositions to an even sampling grid comprises generating a set ofcorrected x, y, and z coordinates of the new corrected position of eachdesired image voxel beneath the refracting surface.
 66. Thenon-transitory computer readable medium according to claim 65, whereininterpolating of new refracted corrected voxel positions to an evensampling grid further comprises producing a set of interpolatedcorrected image brightness values as a function of x, y, and zcoordinates of the even sampling grid and where calculations are appliedafter correcting the raw OCT image dataset for scanning non-linearitiesusing the correction factors derived during calibration.
 67. Thenon-transitory computer readable medium according to claim 46, whereincomputing clinical outputs from corrected image data comprisesgenerating a set of interpolated corrected image brightness values as afunction of x, y, and z coordinates of the even sampling grid.
 68. Thenon-transitory computer readable medium according to claim 67, whereincomputing clinical outputs from corrected image data comprises producinga set of clinical output measures comprising numerical values, tables,graphs, images, and maps.
 69. The non-transitory computer readablemedium according to claim 46, wherein the raw OCT image dataset is of acornea or a lens and wherein computing clinical outputs comprises one ormore of generating thickness maps, generating elevation maps, generatingcurvature maps, computing wavefront aberrations, or calculating an anglebetween an iris and an endothelial surface.
 70. The non-transitorycomputer readable medium according to claim 46, wherein the raw OCTimage dataset is of a cornea and wherein computing clinical outputs fromcorrected image data comprises creating a corneal or lens thickness mapby determining the normal distance between the two surfaces in thevolumetric surface model via an analytical or numerical approach. 71.The non-transitory computer readable medium according to claim 70,wherein a Cartesian coordinate system is used to conduct furthercomputations and a recursive half searching algorithm (“RHSA”) numericalsolution is used to solve for (x, y, z) that exists within a defineddistance to the endothelial surface.